Alignment indication for transcutaneous energy transfer

ABSTRACT

System for transcutaneous energy transfer. An implantable medical device, adapted to be implanted in a patient, has componentry for providing a therapeutic output. The implantable medical device has an internal power source and a secondary coil operatively coupled to the internal power source. An external power source, having a primary coil, provides energy to the implantable medical device when the primary coil of the external power source is placed in proximity of the secondary coil of the implantable medical device and thereby generates a current in the internal power source. An alignment indicator reports the alignment as a function of the current generated in the internal power source with a predetermined value associated with an expected alignment between the primary coil and secondary coil.

RELATED APPLICATION

This application is a division of U.S. patent application Ser. No.11/119,361, filed Apr. 29, 2005, and claims priority therefrom.

FIELD

This invention relates to implantable medical devices and, inparticular, to energy transfer devices, systems and methods forimplantable medical devices.

BACKGROUND

Implantable medical devices for producing a therapeutic result in apatient are well known. Examples of such implantable medical devicesinclude implantable drug infusion pumps, implantable neurostimulators,implantable cardioverters, implantable cardiac pacemakers, implantabledefibrillators and cochlear implants. Of course, it is recognized thatother implantable medical devices are envisioned which utilize energydelivered or transferred from an external device.

A common element in all of these implantable medical devices is the needfor electrical power in the implanted medical device. The implantedmedical device requires electrical power to perform its therapeuticfunction whether it be driving an electrical infusion pump, providing anelectrical neurostimulation pulse or providing an electrical cardiacstimulation pulse. This electrical power is derived from a power source.

Typically, a power source for an implantable medical device can take oneof two forms. The first form utilizes an external power source thattranscutaneously delivers energy via wires or radio frequency energy.Having electrical wires which perforate the skin is disadvantageous due,in part, to the risk of infection. Further, continuously couplingpatients to an external power for therapy is, at least, a largeinconvenience. The second form utilizes single cell batteries as thesource of energy of the implantable medical device. This can beeffective for low power applications, such as pacing devices. However,such single cell batteries usually do not supply the lasting powerrequired to perform new therapies in newer implantable medical devices.In some cases, such as an implantable artificial heart, a single cellbattery might last the patient only a few hours. In other, less extremecases, a single cell unit might expel all or nearly all of its energy inless than a year. This is not desirable due to the need to explant andreimplant the implantable medical device or a portion of the device. Onesolution is for electrical power to be transcutaneously transferredthrough the use of inductive coupling. Such electrical power or energycan optionally be stored in a rechargeable battery. In this form, aninternal power source, such as a battery, can be used for directelectrical power to the implanted medical device. When the battery hasexpended, or nearly expended, its capacity, the battery can be rechargedtranscutaneously, via inductive coupling from an external power sourcetemporarily positioned on the surface of the skin.

Several systems and methods have been used for transcutaneouslyinductively recharging a rechargeable used in an implantable medicaldevice.

PCT Patent Application No. WO 01/83029 A1, Torgerson et al, BatteryRecharge Management For an Implantable Medical Device, (Medtronic, Inc.)discloses an implantable medical device having an implantable powersource such as a rechargeable lithium ion battery. The implantablemedical device includes a recharge module that regulates the rechargingprocess of the implantable power source using closed-loop feedbackcontrol. The recharging module includes a recharge regulator, a rechargemeasurement device monitoring at least one recharge parameter, and arecharge regulation control unit for regulating the recharge energydelivered to the power source in response to the recharge measurementdevice. The recharge module adjusts the energy provided to the powersource to ensure that the power source is being recharged under safelevels.

Transcutaneous energy transfer through the use of inductive couplinginvolves the placement of two coils positioned in close proximity toeach other on opposite sides of the cutaneous boundary. The internalcoil, or secondary coil, is part of or otherwise electrically associatedwith the implanted medical device. The external coil, or primary coil,is associated with the external power source or external charger, orrecharger. The primary coil is driven with an alternating current. Acurrent is induced in the secondary coil through inductive coupling.This current can then be used to power the implanted medical device orto charge, or recharge, an internal power source, or a combination ofthe two.

U.S. Pat. No. 5,713,939, Nedungadi et al, Data Communication System forControl of Transcutaneous Energy Transmission to an Implantable MedicalDevice, discloses a data communication system for control oftranscutaneous energy transmission to an implantable medical device. Theimplantable medical device has rechargeable batteries and a single coilthat is employed both for energy transmission and data telemetry.Control circuitry in the implantable device senses battery voltage andcurrent through the battery, encodes those values by the use ofmultiplexer, and transmits the sensed and encoded values through thecoil to an external energy transmission device. The external deviceincludes a coil that is electromagnetically coupled to the coil in theimplantable device for receiving the encoded signals and fortransmitting energy to the implantable device. The external devicedecodes the transmitted values and transmits those to a controller forcontrolling energy transmission.

U.S. Pat. No. 6,212,431, Hahn et al, Power Transfer Circuit forImplanted Devices, discloses an external power transfer circuit whichcouples ac power having a fixed frequency into an implantable electricalcircuit, e.g., an implantable tissue stimulator, while automaticallymaintaining optimum power transfer conditions. Optimum power transferconditions exist when there is an impedance match between the externaland implanted circuits. The external transfer circuit includes adirectional coupler and an impedance matching circuit. The directionalcoupler senses the forward power being transferred to the implantdevice, as well as the reverse power being reflected form the implantdevice (as a result of an impedance mismatch). The impedance matchingcircuit includes at least one variable element controlled by a controlsignal. The sensed reverse power is used as a feedback signal toautomatically adjust the variable element in the impedance matchingcircuit, and hence the output impedance of the external power transfercircuit, so that it matches the input impedance of the implant device,despite variations that occur in the input impedance of the implantdevice due to variations in implant distance and implant load.

For implanted medical devices, the efficiency at which energy istranscutaneously transferred is crucial. First, the inductive coupling,while inductively inducing a current in the secondary coil, also has atendency to heat surrounding components and tissue. The amount ofheating of surrounding tissue, if excessive, can be deleterious. Sinceheating of surrounding tissue is limited, so also is the amount ofenergy transfer which can be accomplished per unit time. The higher theefficiency of energy transfer, the more energy can be transferred whileat the same time limiting the heating of surrounding components andtissue. Second, it is desirable to limit the amount of time required toachieve a desired charge, or recharge, of an internal power source.While charging, or recharging, is occurring the patient necessarily hasan external encumbrance attached to their body. This attachment mayimpair the patient's mobility and limit the patient's comfort. Thehigher the efficiency of the energy transfer system, the faster thedesired charging, or recharging, can be accomplished limiting theinconvenience to the patient. Third, amount of charging, or recharging,can be limited by the amount of time required for charging, orrecharging. Since the patient is typically inconvenienced during suchcharging, or recharging, there is a practical limit on the amount oftime during which charging, or recharging, should occur. Hence, the sizeof the internal power source can be effectively limited by the amount ofenergy which can be transferred within the amount of charging time. Thehigher the efficiency of the energy transfer system, the greater amountof energy which can be transferred and, hence, the greater the practicalsize of the internal power source. This allows the use of implantablemedical devices having higher power use requirements and providinggreater therapeutic advantage to the patient and/or extends the timebetween charging effectively increasing patient comfort.

Alignment of an external primary coil with the internal secondary coilis important in achieving efficiency in transcutaneous energy transfer.However, it is not always easy for the user to know when the primary andsecondary coils are properly aligned. Often the user must resort totactile information gleaned from the physical package into which theprimary coil is located and a subtle protrusion under the skinapproximately where the implantable medical device has been implanted.However, even perfectly aligning the physical package containing theprimary coil with the protrusion of the implanted medical device may notresult in optimum alignment of the primary and secondary coils. Oftenthe primary coil or the secondary coil, or both, is not centered in thephysical packages within which they are contained. Thus, even perfectalignment of the packages may result in actual misalignment of theprimary and secondary coils.

SUMMARY

Various aspects of the present invention provide a system fortranscutaneous energy transfer, external power system for transcutaneousenergy transfer or a method for indicating alignment between an externalprimary coil and an implanted secondary coil. Alternatively, the presentinvention provides a system for transcutaneous energy transfer having anexternal power source which varies its power output in order to generatea predetermined current in the internal power source, which can be afunction of alignment between the coils, without actually indicatingalignment to a user or other person, e.g., a medical professionalassisting or performing the transcutaneous energy transfer.

In an embodiment, the present invention provides a system fortranscutaneous energy transfer. An implantable medical device, adaptedto be implanted in a patient, has componentry for providing atherapeutic output, an internal power source and a secondary coiloperatively coupled to the internal power source, the implantablemedical device. An external power source, having a primary coil,provides energy to the implantable medical device when the primary coilof the external power source is placed in proximity of the secondarycoil of the implantable medical device and thereby generates a currentin the internal power source. The external power source varies its poweroutput in order to generate a predetermined current in the internalpower source.

In a preferred embodiment, the predetermined current in the internalpower source varies as a function of the voltage of the internal powersource.

In another embodiment, the present invention provides a method oftranscutaneous energy transfer between an external primary coil and aninductively coupled secondary coil of an implanted medical device. Thesecondary coil supplies power to a power source having an internalimpedance. The external primary coil is driven with a charging signal. Acurrent generated in the power by the charging signal is measured. Thecharging signal is varied in order to generate a predetermined currentin the internal power source.

In a preferred embodiment, the method additionally varies thepredetermined current in the internal power source as a function of thevoltage of the internal power source.

In a preferred embodiment, the predetermined current in the internalpower source declines as the voltage of the internal power sourceincreases during a charging cycle.

In a preferred embodiment, the predetermined current in the internalpower source comprises a maximum amount current for charging theinternal power source.

In a preferred embodiment, the predetermined current in the internalpower source declines over time as the internal impedance of theinternal power source increases.

DRAWINGS

FIG. 1 illustrates an implantable medical device implanted in a patient;

FIG. 2 is a block diagram of an implantable medical device;

FIG. 3 is a detailed block diagram of an implantable medical deviceimplanted subcutaneously and an associated external charging device inaccordance with an embodiment of the present invention;

FIG. 4 is a perspective view of an internal antenna associated with animplantable medical device;

FIG. 5 is a side view of the internal antenna of FIG. 4;

FIG. 6 is an exploded perspective view an external antenna andassociated bracket in accordance with an embodiment of the presentinvention;

FIG. 7 is a top view of an external antenna in accordance with anembodiment of the present invention;

FIG. 8 is a perspective view of an external antenna and bracketcombination in accordance with an embodiment of the present invention;

FIG. 9 is a cross-sectional side view of an implantable medical deviceimplanted subcutaneously and an associated bracket for use with anexternal antenna;

FIG. 10 is a cut-away top view of view a primary coil and associatedmagnetic core in accordance with an embodiment of the present invention;

FIG. 11 is a cross-sectional view of the primary coil and associatedmagnetic core of FIG. 10 taken through section line B-B;

FIG. 12 is an exploded view a portion of an external antenna constructedin accordance with an embodiment of the present invention showing themagnetic core and a core cup assembly;

FIG. 13 is block diagram of an external charging unit and an associatedinductively coupled cradle for recharging the external charging unit;

FIG. 14 is a detailed block diagram of the external charging unit ofFIG. 13;

FIG. 15 is a flow chart illustrating a charging process in accordancewith an embodiment of the present invention;

FIG. 16 is a schematic diagram of a dual range temperature sensor;

FIG. 17 is a block diagram of an alignment indicator;

FIG. 18 is a diagram of a display for the alignment indicator of FIG.17; and

FIG. 19 is a flow chart illustrating charging of an implantable medicaldevice.

DESCRIPTION

The entire content of U.S. application Ser. No. 11/119,361, filed Apr.29, 2005, is hereby incorporated by reference.

FIG. 1 shows implantable medical device 16, for example, a drug pump,implanted in patient 18. The implantable medical device 16 is typicallyimplanted by a surgeon in a sterile surgical procedure performed underlocal, regional, or general anesthesia. Before implanting the medicaldevice 16, a catheter 22 is typically implanted with the distal endposition at a desired therapeutic delivery site 23 and the proximal endtunneled under the skin to the location where the medical device 16 isto be implanted. Implantable medical device 16 is generally implantedsubcutaneously at depths, depending upon application and device 16, offrom 1 centimeter (0.4 inches) to 2.5 centimeters (1 inch) where thereis sufficient tissue to support the implanted system. Once medicaldevice 16 is implanted into the patient 18, the incision can be suturedclosed and medical device 16 can begin operation.

Implantable medical device 16 operates to infuse a therapeutic substanceinto patient 18. Implantable medical device 16 can be used for a widevariety of therapies such as pain, spasticity, cancer, and many othermedical conditions.

The therapeutic substance contained in implantable medical device 16 isa substance intended to have a therapeutic effect such as pharmaceuticalcompositions, genetic materials, biologics, and other substances.Pharmaceutical compositions are chemical formulations intended to have atherapeutic effect such as intrathecal antispasmodics, pain medications,chemotherapeutic agents, and the like. Pharmaceutical compositions areoften configured to function in an implanted environment withcharacteristics such as stability at body temperature to retaintherapeutic qualities, concentration to reduce the frequency ofreplenishment, and the like. Genetic materials are substances intendedto have a direct or indirect genetic therapeutic effect such as geneticvectors, genetic regulator elements, genetic structural elements, DNA,and the like. Biologics are substances that are living matter or derivedfrom living matter intended to have a therapeutic effect such as stemcells, platelets, hormones, biologically produced chemicals, and thelike. Other substances may or may not be intended to have a therapeuticeffect and are not easily classified such as saline solution,fluoroscopy agents, disease diagnostic agents and the like. Unlessotherwise noted in the following paragraphs, a drug is synonymous withany therapeutic, diagnostic, or other substance that is delivered by theimplantable infusion device.

Implantable medical device 16 can be any of a number of medical devicessuch as an implantable therapeutic substance delivery device,implantable drug pump, cardiac pacemaker, cardioverter or defibrillator,as examples.

In FIG. 2, implantable medical device 16 has a rechargeable power source24, such as a Lithium ion battery, powering electronics 26 and therapymodule 28 in a conventional manner. Therapy module 28 is coupled topatient 18 through one or more therapy connections 30, alsoconventionally. Rechargeable power source 24, electronics 26 and therapymodule 28 are contained in hermetically sealed housing 32. Secondarycharging coil 34 is attached to the exterior of housing 32. Secondarycharging coil 34 is operatively coupled through electronics 26 torechargeable power source 24. In an alternative embodiment, secondarycharging coil 34 could be contained in housing 32 or could be containedin a separate housing umbilically connected to electronics 26.Electronics 26 help provide control of the charging rate of rechargeablepower source 24 in a conventional manner. Magnetic shield 36 ispositioned between secondary charging coil 34 and housing 32 in order toprotect rechargeable power source 24, electronics 26 and therapy module28 from electromagnetic energy when secondary charging coil 34 isutilized to charge rechargeable power source 24.

Rechargeable power source 24 can be any of a variety power sourcesincluding a chemically based battery or a capacitor. Rechargeable powersource may be a well known lithium ion battery.

FIG. 3 illustrates an alternative embodiment of implantable medicaldevice 16 situated under cutaneous boundary 38. Implantable medicaldevice 16 is similar to the embodiment illustrated in FIG. 2. However,charging regulation module 42 is shown separate from electronics 26controlling therapy module 28. Again, charging regulation and therapycontrol is conventional. Implantable medical device 16 also has internaltelemetry coil 44 configured in conventional manner to communicatethrough external telemetry coil 46 to an external programming device(not shown), charging unit 50 or other device in a conventional mannerin order to both program and control implantable medical device and toexternally obtain information from implantable medical device 16 onceimplantable medical device has been implanted. Internal telemetry coil44, rectangular in shape with dimensions of 1.85 inches (4.7centimeters) by 1.89 inches (4.8 centimeters) constructed from 150 turnsof 43 AWG wire, is sized to be larger than the diameter of secondarycharging coil 34. Secondary coil 34 is constructed with 182 turns of 30AWG wire with an inside diameter of 0.72 inches (1.83 centimeters) andan outside diameter of 1.43 inches (3.63 centimeters) with a height of0.075 inches (0.19 centimeters). Magnetic shield 36 is positionedbetween secondary charging coil 34 and housing 32 and sized to cover thefootprint of secondary charging coil 34.

Internal telemetry coil 44, having a larger diameter than secondary coil34, is not completely covered by magnetic shield 36 allowing implantablemedical device 16 to communicate with the external programming devicewith internal telemetry coil 44 in spite of the presence of magneticshield 36.

Rechargeable power source 24 can be charged while implantable medicaldevice 16 is in place in a patient through the use of external chargingdevice 48. In an embodiment, external charging device 48 consists ofcharging unit 50 and external antenna 52. Charging unit 50 contains theelectronics necessary to drive primary coil 54 with an oscillatingcurrent in order to induce current in secondary coil 34 when primarycoil 54 is placed in the proximity of secondary coil 34. Charging unit50 is operatively coupled to primary coil by cable 56. In an alternativeembodiment, charging unit 50 and antenna 52 may be combined into asingle unit. Antenna 52 may also optionally contain external telemetrycoil 46 which may be operatively coupled to charging unit 50 if it isdesired to communicate to or from implantable medical device 16 withexternal charging device 48. Alternatively, antenna 52 may optionallycontain external telemetry coil 46 which can be operatively coupled toan external programming device, either individually or together withexternal charging unit 48.

As will be explained in more detail below, repositionable magnetic core58 can help to focus electromagnetic energy from primary coil 46 to moreclosely be aligned with secondary coil 34. Also as will be explained inmore detail below, energy absorptive material 60 can help to absorb heatbuild-up in external antenna 52 which will also help allow for a lowertemperature in implantable medical device 16 and/or help lower rechargetimes. Also as will be explained in more detail below, thermallyconductive material 62 is positioned covering at least a portion of thesurface of external antenna 52 which contacts cutaneous boundary 38 ofpatient 18.

As shown in FIG. 4 and FIG. 5, secondary coil 34 and magnetic shield 36are separate from but adjacent to housing 32 encompassing the remainderof implantable medical device 16. Internal antenna 68 is contained in aseparate housing 74 which is attachable to housing 32 so thatimplantable medical device 16 can be implanted by a medical professionalas essentially one unit. Secondary coil 34 is electrically attached tocharging regulation module 42 through leads 82.

In order to achieve efficient inductive coupling between primary coil 54of external antenna 52 and secondary coil 34, it is desirable to placeprimary coil 54 of external antenna 52 as close to secondary coil 34 aspossible. Typically, external antenna 52 is placed directly on cutaneousboundary 38 and, since the location of implantable medical device 16 isfixed, the distance across cutaneous boundary 38 between primary coil 54and secondary coil 34 is minimized as long as external antenna 52 iskept adjacent cutaneous boundary 38.

External antenna 52 is attachable to patient 18 with bracket 84 whencharging rechargeable power source 24. FIG. 6 is an explodedillustration of an embodiment of external antenna 52 attachable tobracket 84. Primary coil 54 is contained in bobbin assembly 86 whichsits in bottom housing 88. Primary coil is connectable to cable 56. Thebottom of external antenna 52 is formed from a thermally conductivematerial 90. Rotating core cup assembly 92 is held in place by tophousing 94. Rotating core cup assembly 92 is rotatable is allowed torotate within external antenna 52. Detents 96 engage detent spring 98 toposition rotatable core cup assembly 92 in one of a plurality of detentpositions. External antenna may be secured together, for example, withscrews (not shown) holding top housing 94 and thermally conductivematerial 90 together.

Bracket 84 is adapted to be attached to the body of patient 18 with abelt (not shown) attachable to bracket 84 with belt loops 102. Ears 104are adapted to mate with tabs 106 in top housing 94 and pivotally secureexternal antenna 52 in bracket 84 when charging is to be accomplished.Bracket 84 has an opening 108 allowing thermally conductive material 90of external antenna 52 to contact the skin of patient 18 when externalantenna 52 is pivotally secured in bracket 84.

As bracket 84 is attached to patient 18 with a belt via belt loops 102,the skin surface of patient 18 is typically not completely flat. Forexample, if implantable medical device 16 is implantable in the bodytorso of patient 18, then the belt attached via belt loops 102 willtypically pass around the torso of patient 18. Since the torso ofpatient 18, and especially the torso of patient 18 near the location ofimplantable medical device 16, bracket 84 may not sit completely flat onpatient 18. This may be especially true as patient 18 moves and thetorso flexes during such movement. Bracket 84 may be conformal andflexible in order to conform to the shape of the body of patient 18.However, bracket 84 may also be rigid enough so that opening 108 inbracket 84 maintains its shape in order to properly receive externalantenna 52. Bracket 84 is preferably constructed of PCABS. To maintainthe proper position of bracket 84 with the skin of patient 18, thesurface of bracket 84 closest to patient 18 contains material 109constructed from a high durometer, e.g., 40 Shore A, or “sticky”material such as a material known under the trade name of “Versaflex”manufactured by GLS Corp. of McHenry, Ill. This will help externalantenna to sit more closely to the skin surface of patient 18 and remainthere during movements of patient 18 throughout the charge or rechargecycle. In addition, external antenna 52 is allowed to pivot by way ofears 104 on tabs 106. Bracket 84 is configured to allow thermallyconductive material 90 to extend through opening 108 and contact theskin surface of patient 18. Allowed pivoting of external antenna 52 and,hence, thermally conductive material 90, permits thermally conductivesurface to sit more closely to the skin surface of patient 18.

FIG. 7 is a partially cut away top view of external antenna 52 inassembled form and attached to cable 56. Rotatable core cup assembly 92is shown located inside of primary coil 54 and positionable in selectedrotated positions via detents 96 and detent spring 98. In FIG. 7,rotatable core cup assembly is positioned between with detent spring 98between detents 96 illustrating that while multiple detent positions areavailable, rotatable core cup assembly can be positioned between detentpositions and, indeed, at any rotated position.

In FIG. 8, the assembly of external antenna 52 with bracket 84 is shownconnected to cable 56. Bracket 84 may be affixed to patient 18 throughbelt loops 102 and then, after bracket 84 has been affixed to patient18, external antenna 52 be attached to bracket 84. Affixing bracket 84to patient 18 first allows for bracket 84 to be used to laterallyposition external antenna close to the position of implantable medicaldevice 16.

Typical prior art positioning systems rely on the external antenna forlateral positioning. The external antenna is moved around on the body ofthe patient 18 until the best lateral position is found. When the bestlateral position is found, the external antenna is removed from the bodyand the bottom of the external antenna (the portion of the externalantenna) contacting the patient's body) is made to be resistant tolateral movement. As an example, one way is to remove a protective linerexposing a sticky surface allowing the external antenna to be relativelyfixed in location. However, the very act of lifting the external antennain order to remove the protective liner and replacing the externalantenna on the body of the patient 18 causes crucial positioninginformation to be lost. There is no guarantee, and in fact it is notlikely, that the external antenna will be replaced in the exact sameposition as the position previously found to be best.

In contrast, bracket 84 of the present invention can be used to roughlyfind the optimum position for external antenna 52. This can be donerelatively easily due to opening 108 in bracket 84. Implantable medicaldevice 16, when implanted, usually leaves an area of the body of patient18 which is not quite as flat as it was before implantation. That is,implantable medical device 16 usually leaves an area of the skin ofpatient 18 which bulges somewhat to accommodate the bulk of implantablemedical device 16. It is relatively easy for patient, medicalprofessional or other person, to place bracket 84 in the general area ofimplantable medical device 16 and move bracket 84 around until the bulgecaused by implantable medical device 16 is most closely centered inopening 108. As bracket 84 is moved laterally, opening 108 tends tonaturally center on the bulge created by implantable medical device 16.Once positioned in this manner, bracket 84 can be secured to the body ofpatient 18 with belt (not shown) attached via belt loops 102. Securingand/or tightening, by pulling the belt tight or snapping a buckle, forexample, can be without removing bracket 84 from the body of patient 16.Thus, bracket 84 can be relatively easily positioned over the generallocation of implantable medical device 16 and secured in that positionwithout being removed from the body of patient 18.

FIG. 9 is cross-sectional view of implantable medical device 16implanted in patient 18 approximately one centimeter under cutaneousboundary 38 creating bulging area 110, an area of the body of patient 18in which the skin of patient 18 is caused to bulge slightly due to theimplantation of implantable medical device 16. Bulging area 110 is anaid to locating the position of external antenna 52 relative tosecondary coil 34. Bracket 84 can be positioned roughly in the areawhere implantable medical device 16 is implanted. Opening 108 in bracket84 can aid in establishing the location of implantable medical device.Bracket 84 can be roughly centered over bulging area 110. After externalantenna 52 is coupled to bracket 84, then primary coil 54 can begenerally centered on implantable medical device 16.

However, secondary coil 34 may not be centered with respect toimplantable medical device 16. This can occur due to a variety ofreasons such as the need for operatively coupling secondary coil 34 tocharging regulation module 42. Connections to make this operativecoupling may require physical space on one side of internal antenna 68which may cause secondary coil 34 not to be centered on implantablemedical device 16. It is also possible that the attachment of internalantenna 68 to housing 32 can cause secondary coil 34 not to be centeredon implantable medical device 16. Regardless of the cause, if secondarycoil 34 is not centered on implantable medical device 16, then centeringbracket 84 on bulging area 110 may not optimally position primary coil54 with respect to secondary coil 34. Any offset in the position ofprimary coil 54 and secondary coil 34 may not result in the mostefficient energy transfer from external antenna 52 to implantablemedical device 16.

A magnetic core 58 is positioned within primary coil 54 in order tofocus energy generated by primary coil 54. Magnetic core 58 attracts themagnetic flux lines generated by primary coil 54. The position ofmagnetic core 58 within primary coil 54 determines the lateral locationof the largest amount of the flux lines generated by primary coil 54.FIGS. 10 and 11 show cut-away top and cross-sectional views of magneticcore 58 used with primary coil 54. Magnetic core 58 is moveable withinprimary coil 54. Lower portion 122 of magnetic core 58 can be rotated toa plurality of positions within primary coil 58 by rotating core cupassembly 92 (see FIG. 12). The travel path of magnetic core 58 can belocked in a plurality of discrete positions. Magnetic core 58 may belocked in four (4) different positions by detents 96 and detent spring98 (see FIG. 6). Magnetic core 58 has an upper planar portion 120 and asmaller lower portion 122.

As magnetic core 58 is repositioned within primary coil 54, the focus ofmagnetic flux generated by primary coil 54 is also repositioned. Asnoted above, external antenna 52 is generally aligned with implantedmedical device 16 using palpatory sensation. Moveable magnetic core 58can then be used to provide a “fine” adjustment to the lateralpositioning of external antenna 52 with respect to secondary coil 34.After bracket 84 has been secured to patient 18, external antenna 52 isattached to bracket 84. Magnetic core 58 is then moved until the bestlateral alignment with secondary coil 34.

Magnetic core 58 is shown positioned within external antenna 52 of FIG.12. Core cup assembly 92 holds magnetic core 58 within the assembly ofexternal antenna 52. Lower portion 122 (not visible in FIG. 12) ofmagnetic core 58 fits into recess 124 of core cup assembly 92 whileupper portion 120 of magnetic core 58 rests upon ledge 126 of core cupassembly 92. Preferably, magnetic core 58 is a ferrite core. Still morepreferably, magnetic core 58 is constructed from MN60LL highperformance, low loss ferrite manufactured by Ceramic Magnetics, Inc.,Fairfield, N.J. Magnetic core 58 has an initial permeability of 6,500and a maximum permeability of 10,500 (typical) with a volume resistivityof 500 ohm-centimeters.

A surface, preferably the top, of magnetic core 58 is lined with anadhesive coated foam 127 and contained in core cup assembly 92. Magneticcore 58 has a tendency to be brittle. Containing magnetic core 58 incore cup assembly assures that even if magnetic core 58 has one or morefractures, magnetic core 58 will still be properly positioned andcontinue to function. Foam 127 also helps to hold magnetic core 58together and minimize gaps between fractured segments of magnetic core58. Further, foam 127 adds mechanical stability to magnetic core 58helping to cushion magnetic core 58 against mechanical impacts, such asfrom dropping external antenna 52 against a hard surface, and helps toprevents audible rattles which may otherwise develop from a fracturedmagnetic core 58.

As shown in FIG. 13, external charging device 48 can be powered eitherdirectly from internal (to charging unit 50) batteries 160 or indirectlyfrom desktop charging device 162. Desktop charging device is connectablevia power cord 164 to a source of AC power, such as a standard readilyavailable wall outlet. Desktop charging device 162 can be configured asa cradle which can receive charging unit 50. Other forms of connectionfrom desktop charging device 162 to a power source, such as by adedicated line cable can also be utilized. Desktop charging device 162can charge and/or recharge batteries 160 in charging unit 50, preferablyby inductive coupling using coil 167 positioned in desktop chargingdevice 162 and coil 168 positioned within charging unit 50. Once chargedand/or recharged, batteries 160 can provide the power through internalcircuitry 168 and cable 56 to external antenna 52. Since charging unit50 may not be coupled directly to the line voltage source of AC power,charging unit 50 may be used with external antenna 52 to transfer powerand/or charge implanted medical device 16 while desktop charging device162 is coupled to a line voltage source of AC power. The inductivecoupling using coil 167 and coil 168 break the possibility of a directconnection between the line voltage source of AC power and externalantenna 52. Batteries 160 also allow charging unit 50 and, hence,external charging device 48, to be used in transferring power and/orcharging of implanted medical device 16 while completely disconnectedfrom either a line voltage source of AC power or desktop charging device162. This, at least in part, allows patient 18 to be ambulatory whiletransferring power and/or charging implanted medical device 16.

FIG. 14 is a block diagram of external charging device 48 controlled bymicroprocessor 212. Transmit block 214 consists of an H-bridge circuitpowered from 12 volt power supply 216. Transmit block 214 drives primarycoil 54 in external antenna 52. H-bridge control signals and timing areprovided conventionally by microprocessor 212. H-bridge circuit intransmit block 214 is used to drive both primary coil 54, used for powertransfer and/or charging, and telemetry antenna 218. Drive selection isdone by electronically controllable switch 220. During power transferand/or charging, H-bridge circuit is driven at 9 kiloHertz. Duringtelemetry, H-bridge circuit is driven at 175 kiloHertz.

Receive block 222 is used only during telemetry, enabled by switch 224,to receive uplink signals from implanted medical device 16. Twelve voltpower supply 216 is a switching regulator supplying power to transmitblock 214 during power transfer and/or charging as well as telemetrydownlink. Nominal input voltage to 12 volt power supply 216 is either7.5 volts from lithium ion batteries 226 or 10 volts from desktopcharging device 162 (FIG. 13).

Current measure block 226 measures current to 12 volt power supply 216.Current measured by current measure block 226 is used in the calculationof power in along with the voltage of batteries 160. As noted above,power in is used in the calculation of efficiency of power transferand/or charging efficiency to determine, in part, the best location ofexternal antenna 52 and/or rotating core cup assembly 92.

Rotating core cup assembly 92 is rotated in external antenna 52 forbetter lateral alignment of primary coil 54 and secondary coil 34. Afeedback mechanism is used to determine the best rotation of core cupassembly 92. External charging device 48 can determine whether thecurrent position of rotating core cup assembly 92 is optimally alignedfor energy transfer and/or charging. External charging device 48measures the power out of external charging device 48 divided by thepower into external charging device 48. This calculation is a measure ofthe efficiency of external charging device 48. The power out is gaugedby the power induced in implantable medical device 16 and is determinedby multiplying the voltage of rechargeable power source 24 by thecharging current in implantable medical device 16. These values areobtained by telemetry from implanted medical device 16. The power in isgauged by the power generated by charging unit 50 and is determined bymultiplying the voltage of the internal voltage of charging unit 50,e.g., the voltage of a battery or batteries internal to charging unit50, by the current driving external antenna 52, in particular I-rechargefrom Imeasure unit 226.

The ratio of power out divided by power in can be scaled displayed topatient 18, or a medical professional or other person adjustingrotatable core cup assembly 92 or positioning external antenna 52. Forexample, the available efficiency can be divided into separate rangesand displayed as a bar or as a series of lights. The separate ranges canbe linearly divided or can be logarithmic, for example.

Using efficiency as a measure of effective coupling and, hence, as ameasure of proper location of external antenna 52 and rotatable core cupassembly 92 works not only at high charging or power transfer levels butalso at reduced charging levels, as for example, when charging atreduced levels toward the end or beginning of a charging cycle.

If, after patient 18 or other person has moved rotatable core cupassembly 92 through all of the range of positions on external antenna 52and can not achieve an acceptable efficiency level, patient 18 or otherperson can remove external antenna 52 from bracket 84, realign bracket84 with bulging area 110, reattach external antenna 52 to bracket 84 andrestart the alignment and coupling efficiency process.

FIG. 15 is a flow chart illustrating an exemplary charging process usingexternal antenna 52. The process starts [block 126] and a chargingsession begins [block 128] with a test [block 130]. The charging systemperforms start-up checks [block 132]. If the start-up checks are notperformed successfully, the actions taken in Table 1 are performed.

TABLE 1 Check Screen/Message System Errors: e.g., stuck key System ErrorExternal Charger Battery Status Recharge Complete Battery Low RechargeExternal Charger External Charger Connected to External Recharge inProcess Icon Antenna Antenna Disconnect Connect Antenna

If the start-up checks are successful, telemetry with implantablemedical device 16 is checked [block 134]. If telemetry is unsuccessful,the error messages indicated in Table 2 are generated.

TABLE 2 Failure Screen/Message Poor Communication Reposition AntennaExternal Charger Error Code Response Call Manufacturer CommunicationError Communication Error External Charger Fault Call ManufacturerAntenna Disconnect Connect Antenna Antenna Failure Antenna Failure Icon

If telemetry checks are successful, external charging device 48 is ableto monitor [block 136] charging status. Monitoring charging status canincludes providing feedback to an operator to help determine the bestrotational position of core cup assembly 92.

Charge events are checked [block 138]. If no charge events are noted,the actions indicated in Table 3 are executed.

TABLE 3 Event Screen/Message Telemetry Failure (See Messages From Table2) Implantable Medical Device Battery Device Battery Low Low ExternalCharger Battery Low Charger Battery Low External Charger BatteryDepleted Recharge Charger External Charger Recharge Complete ExternalCharger Recharge Complete Implantable Medical Device Will Not RechargeDevice Provide Therapeutic Result Until Recharged: Therapy Unavailable/Sleep Mode Antenna Disconnect Connect Antenna

If a charge event occurs, then the process checks to determine ifcharging is complete [block 140]. Once charging is complete, the processterminates [block 142].

As energy is transferred from primary coil 54 of external antenna 52 tosecondary coil 34 of implantable medical device 16, heat may also begenerated in implantable medical device 16 in surrounding tissue ofpatient 18. Such heat build-up in tissue of patient 18, beyond certainlimits, is undesirable and should be limited as acceptable values.Generally, it is preferable to limit the temperature of external antenna52 to not more than forty-one degrees Centigrade (41° C.) and to limitthe temperature of implanted medical device 16 and the skin of patient18 to thirty-nine degrees Centigrade (39° C.). In order to ensure thatimplantable medical device 16 is less than the upper limit ofthirty-nine degrees Centigrade (39° C.), the actual temperature ofexternal antenna 52 may be less than thirty-nine degrees Centigrade (39°C.). In general, the temperature of external antenna 52 may bemaintained to be less than or equal to the desired maximum temperatureof implanted medical device 16. While the temperature limits discussedabove are anticipated under current conditions and regulations, it isrecognized and understood that conditions and regulations may change orbe different in different circumstances. Accordingly, the actualtemperatures and temperature limits may change. Such temperature limitsmay be under software control in charging unit 50 so that any suchtemperatures or temperature limits can be modified to fit the thencurrent circumstances.

Magnetic shield 36 serves to at least partially protect the portion ofimplantable medical device 16 contained within titanium housing 32 fromthe effects of energy transfer from external charging device 48 producedthrough inductive coupling from primary coil 54. Magnetic shield 36 isconstructed of Metglas magnetic alloy 2714A (cobalt-based) manufacturedby Honeywell International, Conway, S.C. Magnetic shield 36 ispositioned between secondary coil 34 and housing 32 of implantablemedical device 16 with secondary coil 34 facing cutaneous boundary 38.Magnetic shield does not interfere with the operation of secondary coil34 because magnetic shield 36 is positioned away from primary coil 54.Also, magnetic shield does not interfere with telemetry betweenimplantable medical device 16 and an external programmer becausemagnetic shield 36 is smaller than internal telemetry coil 44. That is,internal telemetry coil 44 lies outside of magnetic shield 36.

However, the material of magnetic shield 36 substantially limits theelectromagnetic energy induced by primary coil 54 from penetratingbeyond magnetic shield. Electromagnetic waves induced by primary coil 54that reach titanium housing 32 will tend to be absorbed by titaniumhousing 54 and its components and will tend to cause the temperature oftitanium housing 54 to rise. As the temperature of titanium housing 54rises, such temperature increase will be disadvantageously transferredto the surrounding tissue of patient 18. However, any electromagneticwaves which are prevented from reaching titanium housing 32 will notcause such a temperature rise.

Thermally conductive material 62 of external antenna 52 is positioned tocontact the skin of patient 18 when external antenna 52 is placed forenergy transfer, or charging, of implanted medical device 16. Thermallyconductive material 62 tends to spread any heat generated at the skinsurface and spread any such heat over a larger area. Thermallyconductive material 62 tends to make the temperature of the skin surfacemore uniform than would otherwise be the case. Uniformity of temperaturewill tend to limit the maximum temperature of any particular spot on theskin surface. The skin itself is a pretty good conductor of heat andinitially spreading any heat generated over a larger area of the skinwill further assist the skin in dissipating any heat build-up on tosurrounding tissue and further limit the maximum temperature of anyparticular location on the surface of the skin.

Thermally conductive material 62 is molded into the surface of externalantenna 52 which will contact the skin surface of patient 18 whenexternal antenna 52 provides energy transfer to implanted medical device16. Since thermally conductive material 62 should pass electromagneticenergy from primary coil 54, thermally conductive material 62 should beconstructed from a non-magnetic material. It is desirable that thermallyconductive material 62 have a thermal conductivity of approximately 5.62BTU inch/hour feet² degrees Fahrenheit (0.81 W/meters degrees Kelvin).Thermally conductive material may be constructed from a proprietarycomposite of approximately forty percent (40%) graphite, seven percent(7%) glass in RTP 199×103410 A polypropylene, manufactured by RTPCompany, Winona, Minn. Thermally conductive material may not beelectrically conductive in order to reduce eddy currents. Thermallyconductive material may have a volume resistivity of approximately 10³ohm-centimeters and a surface resistivity of 10⁵ ohms per square.

Energy absorptive material 62 can be placed in and/or around primarycoil 54 of external antenna 52 in order to absorb some of the energygenerated by primary coil 54. Energy absorptive material 62 may fill inan otherwise empty space of rotating core cup assembly 92. Heatgenerated by energy produced by primary coil 54 which is not effectivelyinductively coupled to secondary coil 34 will tend to cause atemperature rise in other components of external antenna 52 and,possibly, the skin of patient 18. At least a portion of this temperaturerise can be blocked through the use of energy absorptive material 62.Energy absorptive material 62 is chosen to absorb heat build-up insurrounding components and tend to limit further temperature increases.Preferably, energy absorptive material 62 is selected to be materialwhich undergoes a state change at temperatures which are likely to beencountered as the temperature of surrounding components rises duringenergy transfer, e.g., charging, using external antenna 52.

If it is a goal to limit the temperature of the skin of patient 18 tothirty-nine degrees Centigrade (39° C.), it is desirable to use ofenergy absorptive material 62 which has a state change at or near thetemperature limit. In this example, the use of an energy absorptivematerial 62 having a state change in temperature area just belowthirty-nine degrees Centigrade (39° C.), preferably in the range ofthirty-five degrees Centigrade (35° C.) to thirty-eight degreesCentigrade (38° C.), can help limit the rise in the temperature of theskin of patient 18 to no more than the desired limit, in this example,thirty-nine degrees (39° C.).

As the temperature of surrounding components of external antenna 52 riseto a temperature which is just below the temperature at which energyabsorptive material 62 changes state, at least a portion of further heatenergy generated by primary coil 54 and surrounding components ofexternal antenna 52 will go toward providing the energy necessary forenergy absorptive material 62 to change state. As energy absorptivematerial 62 is in the process of changing state, its temperature is notincreasing. Therefore, during the state change of energy absorptivematerial 62, energy absorptive material 62 is serving to at leastpartially limit a further rise in the temperature of components ofexternal antenna 52. As the state change temperature of energyabsorptive material has been preferably selected to be near or justbelow the temperature limit of the skin of patient 18, energy absorptivematerial 62 will tend to limit the temperature components of externalantenna 52 from reaching the temperature limit and, hence, will alsotend to limit the temperature of the skin of patient 18 from reachingthe maximum desired temperature limit.

Energy absorptive material 62 may be constructed from wax and, inparticular, a wax which has change of state temperature of approximatelythe maximum temperature at which external antenna 52 is desired toreach, such as thirty-eight (38) or thirty-nine (39) degrees Centigrade.Thus, the wax material of which energy absorptive material isconstructed may melt at that temperature.

Inductive coupling between primary coil 54 of external antenna 52 andsecondary coil of implantable medical device 16 is accomplished at adrive, or carrier, frequency, f_(carrier), in the range of from eight(8) to twelve (12) kiloHertz. The carrier frequency, f_(carrier), ofexternal antenna 54 is approximately nine (9) kiloHertz unloaded.

However, the inductive coupling between primary coil 54 of externalantenna 52 and secondary coil 34 of implantable medical device isdependent upon the mutual inductance between the devices. The mutualinductance depends upon a number of variables. Primary coil 54 ispreferably made from a coil of wire that has an inductance L and aseries or parallel tuned capacitance C. The values of both inductance Land capacitance C are fixed. For instance, if the desired drivefrequency, f_(carrier), of the energy transfer system was to be 1megaHertz and external antenna 52 had an independence of one microHenry,capacitance would be added so that the resonant frequency of the energytransfer system would equal that of the drive frequency, f_(carrier).The total capacitance added can be found using the equation f_(resonate)equals one divided by two times pi (π) times the square root of L timesC where L is the inductance of the energy transfer system. In thisexample, the value of capacitance C required to tune external antenna 52to resonate at the carrier frequency of 1 megaHertz is calculated asapproximately 25 nanofarads.

However, when the electrical properties of external antenna 52 change,either by the reflected environment or due to a physical distortion orchange in the composition of the external antenna 52, the inductance, L,may be altered. The inductance, L, can be altered because it is made upof two separate parts. The first part is the self-inductance, L_(self),of external antenna 52 at f_(carrier). The second part is the mutualinductance, L_(mutual), which is a measure of the change in currentdriving external antenna 52 and the magnetic effect, or “loading”, whichthe environment has on external antenna 52. When the electricalcharacteristics of the environment of external antenna 52 change,L_(self) remains constant while L_(mutual) varies. The effect of achange in the overall inductance, whether that change is from L_(self)or from L_(mutual), is a change in the resonant frequency, f_(resonate).Since C was chosen in order to have the resonant frequency,f_(resonate), match the drive frequency, f_(carrier), in order toincrease the efficiency of energy transfer from primary coil 54 ofexternal antenna 52 to secondary coil 34, a change in either or canresult in the resonant frequency, f_(resonate), being mismatched withthe drive frequency, f_(carrier). The result can be a less than optimumefficiency of energy transfer to implantable medical device 16.

As the drive frequency, f_(carrier), varies with respect to the resonantfrequency, f_(resonate), apparent impedance of the energy transfersystem, as seen by primary coil 54, will vary. The apparent impedancewill be at a minimum when the drive frequency, f_(carrier), exactlymatches the resonant frequency, f_(resonate). Any mismatch of the drivefrequency, f_(carrier), from the resonant frequency, will cause theimpedance to increase. Maximum efficiency occurs when the drivefrequency, f_(carrier), matches the resonant frequency, f_(resonate).

As the impedance of the energy transfer system varies, so does thecurrent driving primary coil 54. As the impedance of the energy transfersystem increases, the current driving primary coil 54 will decreasessince the voltage being applied to primary coil 54 remains relativelyconstant. Similarly, the current driving primary coil 54 will increaseas the impedance of the energy transfer system decreases. It can be seenthen that point of maximum current driving primary coil 54 will be at amaximum when the impedance of the energy transfer system is at aminimum, when the resonant frequency, f_(resonate), matches the drivefrequency, f_(carrier), and when maximum efficiency occurs.

The impedance of the energy transfer system can be monitored since thecurrent driving primary coil 54 varies as a function of drive frequency,f_(carrier). The drive frequency can be varied and the current drivingprimary coil can be measured to determine the point at which theimpedance of the energy transfer system is at a minimum, the resonantfrequency, f_(resonate), matches the drive frequency, f_(carrier), andwhen maximum efficiency occurs.

Instead of holding the drive frequency, f_(carrier), constant for anominal resonant frequency, f_(resonate), the drive frequency,f_(carrier), may be varied until the current driving primary coil 54 isat a maximum. This is not only the point at which the impedance of theenergy transfer system is at a minimum but also the point at whichmaximum efficiency occurs.

Maximum efficiency is not as important in systems, such as telemetrysystems, which are utilized in a static environment or for relativelyshort periods of time. In a static environment, the resonant frequency,f_(resonate), may be relatively invariable. Further, efficiency in notterribly important when energy or information transfer occurs over arelatively short period of time.

However, transcutaneous energy transfer systems can be utilized overextended periods of time, either to power the implanted medical device16 over an extended period of time or to charge a replenishable powersupply within implanted medical device 16. Depending upon capacity ofthe replenishable power supply and the efficiency of energy transfer,charging unit 50 can be utilized for hours and typically can be used aspatient 18 rests or over night as patient 18 sleeps. Further, over theextended period of time in which charging unit 50 is utilized, externalantenna 52 is affixed to the body of patient 18. As patient 18 attemptsto continue a normal routine, such as by making normal movement or bysleeping, during energy transfer, it is difficult to maintain externalantenna 52 in a completely fixed position relative to secondary coil 34.Movement of external antenna 52 with respect to secondary coil 34 canresult in a change in mutual inductance, L_(mutual), a change inimpedance and a change in the resonant frequency, f_(resonate). Further,any change in spatial positioning of the energy transfer system with anyexternal conductive object, any change in the characteristics ofexternal antenna 52, such as by fractures in magnetic core 58, forexample, a change in the charge level of rechargeable power source 24 ofimplantable medical device 16 or a change in the power level of chargingunit 50, all can result in a change of mutual inductance, L_(mutual).

Drive frequency, f_(carrier), may be varied, not only initially duringthe commencement of energy transfer, e.g., charging, but also duringenergy transfer by varying the drive frequency, f_(carrier), in order tomatch the drive frequency, f_(carrier), with the resonant frequency,f_(resonate), and, hence, maintaining a relatively high efficiency ofenergy transfer. As an example, drive frequency, f_(carrier), can beconstantly updated to seek resonant frequency, f_(resonate), or drivefrequency, f_(carrier), can be periodically updated, perhaps every fewminutes or every hour as desired. Such relatively high efficiency inenergy transfer will reduce the amount of time charging unit 50 willneed to be operated, for a given amount of energy transfer, e.g., agiven amount of battery charge. A reduced energy transfer, or charging,time can result in a decrease in the amount of heating of implantedmedical device 16 and surrounding tissue of patient 18.

External charging device 48 may incorporate temperature sensor 87 inexternal antenna 52 and control circuitry in charging unit 50 which canensure that external antenna 52 does not exceed acceptable temperatures,generally a maximum of thirty-eight degrees Centigrade (38° C.),preferably 38.3 degrees Centigrade. Temperature sensor 87 in externalantenna 52 can be used to determine the temperature of external antenna52. Temperature sensor 87 can be positioned in close proximity tothermally conductive material 62 in order to obtain reasonably accurateinformation on the temperature of the external surface of externalantenna 52 contacting patient 18. Preferably, temperature sensor 87 isaffixed to thermally conductive material 62 with a thermally conductiveadhesive. Thermally conductive material 62 smoothes out any temperaturesdifferences which otherwise might occur on the surface of externalantenna 52 contacting patient 18. Positioning temperature sensor 87 inthe proximity or touching thermally conductive material 62 enables anaccurate measurement of the contact temperature.

Control circuitry using the output from temperature sensor 87 can thenlimit the energy transfer process in order to limit the temperaturewhich external antenna 52 imparts to patient 18. As temperature sensor87 approaches or reaches preset limits, control circuitry can takeappropriate action such as limiting the amount of energy transferred,e.g., by limiting the current driving primary coil 54, or limiting thetime during which energy is transferred, e.g., by curtailing energytransfer or by switching energy transfer on and off to provide an energytransfer duty cycle of less than one hundred percent.

When the temperature sensed by the temperature sensor is well belowpreset temperature limits, it may be acceptable to report thetemperature with relatively less precision. As an example, if thetemperature sensed by temperature sensor 87 is more than two degreesCentigrade (2° C.) away from a preset limit of thirty-eight degreesCentigrade (38° C.), it may be acceptable to know the temperature withan accuracy of three degrees Centigrade (3° C.).

However, when the temperature of external antenna 52 approaches towithin two degrees Centigrade (2° C.), it may be desirable to know thetemperature with a much greater accuracy, for example, an accuracy ofwithin one tenth of one degree Centigrade (0.1° C.).

It is generally difficult, however, to produce a temperature which has ahigh degree of accuracy over a very broad temperature range. While atemperature sensor can easily be produced to provide a resolution withinone-tenth of one degree Centigrade (0.1° C.) over a relatively narrowrange temperatures, it can be difficult to produce a temperature sensorproviding such a resolution over a broad range of temperatures.

A dual range temperature sensor has a first, broad, less accurate rangeof measurement from thirty degrees Centigrade (30° C.) to forty-twodegrees Centigrade (42° C.) having an accuracy within three degreesCentigrade (3° C.). Further, this temperature sensor has a second,narrow, more accurate range of measurement over four degrees Centigrade(4° C.), from thirty-six degrees Centigrade (36° C.) to forty-twodegrees Centigrade (42° C.), having an accuracy within one-tenth of onedegree Centigrade (0.1° C.).

FIG. 16 illustrates a dual range temperature sensor utilizingtemperature sensor 87. Temperature sensor 87, located in externalantenna 52, is coupled to amplifier 170 which has been pre-calibrated tooperate only in the range of from thirty degrees Centigrade (30° C.) toforty-two degrees Centigrade (42° C.). Components of amplifier 170 havean accuracy reflecting a temperature within one-tenth of one degreeCentigrade (0.1° C.). The analog output of amplifier 170 is sent toanalog-to-digital converter 172 producing a digital output 173 having anaccuracy of one-tenth of one degree Centigrade (0.1° C.). The analogoutput of amplifier 170 could be also sent to comparator 174 whichcompares the analog output against a known reference voltage 176 whichis set to at a predetermined level to produce a positive output 178 whentemperature sensor 87 reflects a temperature of 38.3 degrees Centigrade,the maximum temperature permitted for external antenna 52. Control logicin charging unit 50 can then take appropriate action to limit furthertemperature increases such as by ceasing or limiting further energytransfer and/or charging. Temperature sensor 87 is also coupled toamplifier 182. Components of amplifier 182 have an accuracy reflecting atemperature within three degrees Centigrade (3° C.), much less accuracythan amplifier 170, but amplifier 182 can operate over the much largertemperature range of zero degrees Centigrade (0° C.) to forty-fivedegrees Centigrade (45° C.). The output of amplifier 182 is sent toanalog-to-digital converter 184 producing a digital output 186 having anaccuracy of three degrees Centigrade (3° C.).

FIG. 17 illustrates a functional block diagram of a preferred embodimentof charging unit 50 incorporating alignment indicator 150 and display152. During the process in which rechargeable power source 24 is chargedthrough the use of charging unit 50, it is desirable to obtain theefficiency in power transfer from external charging unit 50 andrechargeable power source 24. Typically, efficiency of energy transferwill be greatest when primary coil 54 of charging unit 50 istranscutaneously optimally aligned with secondary coil 34 of implantablemedical device 16. The subcutaneous position of implantable medicaldevice can often be discerned by a tell-tale bulge in cutaneous 38 andsuch bulge can be used as a guide in placement of primary coil 54.However, secondary coil 34 may not be centered with respect toimplantable medical device 16 and/or such bulge. Similarly, primary coilmay be centered in charging unit 50 or external antenna 52. Thus, it canbe difficult to determine that optimum position for placement of primarycoil 54 by using the physical housing of charging unit 50 or externalantenna 52 in conjunction with a bulge created by the implanted device.

In a preferred embodiment, alignment indicator 150 functionally providesactive feedback to patient 18, or other person, responsible forpositioning primary coil 54 during charging of rechargeable power source24. Alignment may provide sensory feedback, e.g., audible, visual ortactile. In a preferred embodiment, display 152 is utilized to alert auser positioning primary coil 54. In a preferred embodiment, display 152is a series of lights forming a bar graph indicative of a degree ofefficiency of energy transfer, i.e., alignment.

A preferred embodiment of display 152 is represent by bar graph 154,illustrated in FIG. 18. Bar graph 154 consists of eight (8) lights,namely light 188, light 190, light 192, light 194, light 196, light 196,light 198, light 200 and light 202, arranged in a row such that when oneor more lights are lit starting at one end of bar graph 154, the visualappearance of the series of lights resembles a bar graph. In the exampleillustrated in FIG. 18, five (5) of the lights of bar graph 154 areilluminated, namely light 188, light 190, light 192, light 194, light196, giving a visual indication of the alignment of primary coil 54 tosecondary coil 34. Lights 198, 200 and 202 are not illuminated since thedegree of alignment required to achieve such an indication is notpresent. It is to be recognized and understood that the representationof five (5) lights lit in bar graph 154 is merely example and othernumbers of lights, including no lights or all lights, being lit could berepresentative of a degree of alignment.

In a preferred embodiment, Table 4 provides an illustration of how thenumber of lights, representing a bar, may be lit depending upon afunction of the current associated with, or through, secondary coil 34.In this example, the calculated efficiency of energy transfer is used.Efficiency of energy transfer in this embodiment is defined as describedabove with reference to FIG. 14. Table 4 also provides an indication ofan approximate amount of time that would be required for completerecharging of rechargeable power source 24 at the indicated alignment.

TABLE 4 Efficiency Number of Lights Current Time Zero 5.6% 0 14 mA 18hours 5.6% 6.4% 1 16 mA 16 hours 6.4% 7.2% 2 18 mA 14 hours 7.2% 8.0% 320 mA 12 hours 8.0% 10.0% 4 25 mA 10 hours 10.0% 12.4% 5 31 mA 8 hours12.4% 14.8% 6 37 mA 7 hours 14.8% 17.5% 7 44 mA 6 hours 17.5% maximum8 >44 mA  <6 hours

While display 152 has been illustrated and described as a bar graph 154represented by a series of lights, it is to be recognized and understoodthat other forms of display are contemplated, including, but not limitedto, LED displays, LCD displays, plasma displays, numeric oralpha-numeric displays, and the like. Any form of visual representationof the amount of current associated with secondary coil 34 and/or theefficiency of energy transfer can suffice.

It is also to be recognized and understood, while the indication ofalignment has been described and illustrated as a visual indication,that other forms of indication are also contemplated. For example, theindication could be audible, represented by a tone that changes volumeand/or pitch as alignment changes. Further, such an audible indicationcould be verbal. In addition, it is contemplated that tactileindications are contemplated. For example, a vibration that varies inintensity and/or frequency could be used to indicate change inalignment. The indication of alignment could be anything that can bedetected to the sensory perception of a person.

In a preferred embodiment, the alignment indication is based upon theamount of current actually flowing through rechargeable power source 24.It is to be recognized and understood, however, that it is not necessarythat the current measured actually be the current passing throughrechargeable power source 24. Alternatively, an alignment measurementmay be made by measuring a value, e.g., current or voltage, associatedwith, e.g., proportional to, the current passing through rechargeablepower source 24. It is significant, however, that the currentmeasurement be taken on the implant side of the charging system so thatthe measurement of alignment provides an actual indication of alignmentand not a presumed alignment based upon an external measurement.

As rechargeable power source 24 is charged and the voltage acrossrechargeable power source 24 approaches its fully recharged value, itbecomes more difficult to ascertain when alignment is obtained. This isdue, in part, to relatively small changes in the values of the chargingcurrent as a fully charged state is approached. Partly, because of thisreason and partly because current levels and power levels are lower as afully charged state is approached, it may be desirable to discontinuealignment indication following reaching a predetermined point in thecharging cycle. For example, the alignment indication system could ceaseto report alignment when the voltage across rechargeable power source 24reaches ninety percent (90%) of expected fully charged state ofrechargeable power source 24.

Operation of charging unit 50, including alignment indicator 150, isillustrated in the flow chart of FIG. 19. First, charging unit 50determines (310) whether external antenna 52 is over the temperaturelimit set for charging operation. This temperature limit can helpprevent patient 18 from being exposed to temperatures that are higherthan desired. If external antenna 52 of charging unit 50 is overtemperature, an alert condition is indicated (311). If external antennais not over the temperature limit, charging unit 50 then checks (312)for a status problem with charging unit 50. If a status problem isfound, an alert condition is indicated (311).

If a status problem is not found, charging unit 50 initially charges(314) rechargeable power source 24 of implantable medical device 16 for5.5 seconds. Charging unit 50 then stops charging and waits (316) onesecond to check for reception of a telemetry signal from implantablemedical device 16. An example of information sent from implantablemedical device 16 to charging unit 50 via telemetry can include thevalue of a current associated with secondary coil 34, e.g., the value ofthe current flowing through secondary coil 34. If no telemetry signal isdetected, an alert condition is indicated (311). If telemetry isreceived, charging unit 50 then checks (318) for a status problem withimplantable medical device 16. If a status problem is detected, an alertcondition is indicated (311).

If no status problem exists, charging unit 50 checks (322) to determineif the temperature is too high. Again, a temperature exceedingpredetermined limits is not advantageous. If an over temperaturecondition is detected, charging is stopped and a status indication isdisplayed until the temperature drops below a predetermined level.

If no over temperature condition exists, charging unit 50 checks (328)to determine if the voltage across rechargeable power source 24 is overa voltage at which the charging rate should begin to decrease, e.g.,4.05 volts. If the voltage across rechargeable power 24 is greater than4.05 volts, then charging unit 50 begins to taper charging power (330).

If the voltage across rechargeable power source 24 is not over 4.05volts, charging unit 50 checks (332) to determine whether the chargingcurrent through rechargeable power source 24 is over a current rate thatis not desirable, e.g., 50 milliamperes. If the charging current is over50 milliamperes, then the charging power level is decreased (334) by anappropriate, e.g., by 35 milliwatts.

If the charging current is not over 50 milliamperes, charging unit 50checks (336) to determine if the charging power level is less thanappropriate amount, e.g., 925 milliwatts. If the power level is lessthan 925 milliwatts, the charging power level is increased (338) by 35milliwatts, up to a maximum of 925 milliwatts.

If the charge current is below (340) five (5) milliamperes, thencharging unit 50 stops (342) charging and indicates that charging iscomplete, e.g., by lighting the charging complete indicator light.

If not, charging unit 50 then charges (314) rechargeable power sourcefor one (1) minute and then conducts the aforementioned tests, checksand actions as performed after the initial 5.5 second charge.

Thus, embodiments of the alignment indication for transcutaneous energytransfer are disclosed. One skilled in the art will appreciate that thepresent invention can be practiced with embodiments other than thosedisclosed. The disclosed embodiments are presented for purposes ofillustration and not limitation, and the present invention is limitedonly by the claims that follow.

1. A system for transcutaneous energy transfer, comprising: an implantable medical device having componentry for providing a therapeutic output, said implantable medical device having an internal power source and a secondary coil operatively coupled to said internal power source, said implantable medical device adapted to be implanted in a patient; and an external power source having a primary coil, said external power source providing energy to said implantable medical device when said primary coil of said external power source is placed in proximity of said secondary coil of said implantable medical device and thereby generating a current in said internal power source; wherein said external power source automatically varies its power output in order to generate a predetermined current in said internal power source.
 2. The system as in claim 1 wherein said predetermined current in said internal power source varies as a function of said voltage of said internal power source.
 3. The system as in claim 2 wherein said predetermined current in said internal power source declines as said voltage of said internal power source increases during a charging cycle.
 4. The system as in claim 1 wherein said predetermined current in said internal power source comprises a maximum amount current for charging said internal power source.
 5. The system as in claim 4 wherein said predetermined current in said internal power source declines over time as said internal impedance of the said internal power source increases.
 6. The system as in claim 1 wherein said external power source reduces its power output if a current in said internal power source is greater than said predetermined current.
 7. The system as in claim 1 wherein said external power source increases its power output if a current in said internal power source is less than said predetermined current.
 8. The system as in claim 7 wherein said external power source limits its power output to a predetermined power amount.
 9. The system as in claim 1 wherein said external power source terminates its power output if a current in said internal power source is below a minimum amount.
 10. A method of transcutaneous energy transfer between an external primary coil and an inductively coupled secondary coil of an implanted medical device, said external primary coil being operatively coupled to a charging unit, said secondary coil supplying power to a power source having an internal impedance, comprising the steps of: driving said external primary coil with a charging signal from said charging unit; measuring a current generated in said power source by said charging signal; and said charging unit automatically varying said charging signal in order to generate a predetermined current in said internal power source.
 11. The method as in claim 10 further comprising the step of varying said predetermined current in said internal power source as a function of said voltage of said internal power source.
 12. The method as in claim 11 wherein said predetermined current in said internal power source declines as said voltage of said internal power source increases during a charging cycle.
 13. The method as in claim 10 wherein said predetermined current in said internal power source comprises a maximum amount current for charging said internal power source.
 14. The method as in claim 13 wherein said predetermined current in said internal power source declines over time as said internal impedance of the said internal power source increases.
 15. The method as in claim 10 wherein said external power source reduces its power output if a current in said internal power source is greater than said predetermined current.
 16. The method as in claim 10 wherein said external power source increases its power output if a current in said internal power source is less than said predetermined current.
 17. The method as in claim 16 wherein said external power source limits its power output to a predetermined power amount.
 18. The method as in claim 10 wherein said external power source terminates its power output if a current in said internal power source is below a minimum amount.
 19. An external power source for use with an implantable medical device adapted to be implanted in a patient and having componentry for providing a therapeutic output, an internal power source and a secondary coil operatively coupled to said internal power source, comprising: an external power unit; and a primary coil, operatively coupled to said external power unit; said external power source providing energy to said implantable medical device when said primary coil is placed in proximity of said secondary coil of said implantable medical device and thereby generating a current in said internal power source; wherein said external power source automatically varies its power output in order to generate a predetermined current in said internal power source.
 20. The external power source as in claim 19 wherein said external power unit reduces its power output if a current in said internal power source is greater than said predetermined current. 